The present invention is directed to nuclear magnetic resonance imaging and spectroscopy systems and, more particularly, to a novel digital interface subsystem for an imaging and spectroscopic nuclear magnetic resonance (NMR) system. The nuclear magnetic resonance phenomenon occurs in atomic nuclei having an odd number of protons and/or neutrons. Each such nucleus has a net magnetic moment such that when placed in a static homogeneous magnetic field, denoted B.sub.0, a greater number of the involved nuclei become aligned with the B.sub.0 field to create a net magnetization, denoted M, in the direction of the B.sub.0 field. The net magnetization M is the summation of the individual nuclear magnetic moments. Because a nuclear magnetic moment is the result of a nuclear spin, the terms "nuclear moment" and "nuclear spin" are generally used synonymously in the art.
Under the influence of the static homogeneous magnetic field B.sub.0, the nuclei precess, or rotate, about the axis of the B.sub.0 field and hence the net magnetization M is aligned with the B.sub.0 field axis. The rate, or frequency, at which the nuclei precess is dependent upon the strength of the total magnetic field applied to a particular nucleus, and upon the characteristics of the nuclei specie being subjected to the total magnetic field. The angular frequency of precession, .omega., is defined as the Larmor frequency, in accordance with the equation: .omega.=.gamma.B.sub.0, wherein .gamma. is the gyromagnetic ratio (and is constant for each nucleus type) and B.sub.0 is the strength of the total applied magnetic field to the particular nucleus. Thus, the frequency at which the nuclei precess is primarily dependent upon the strength of the total magnetic field B.sub.0 ; the Larmor frequency increases with increasing total magnetic field strength.
A precessing nucleus is capable of resonantly absorbing electromagnetic energy. The frequency of the electromagnetic energy needed to induce resonance is the same Larmor frequency as the precession frequency .omega.. During the application of electromagnetic energy, typically by a pulse of radio-frequency (RF) energy, the net magnetization M precesses further away from the B.sub.0 field axis (arbitrarily assumed to be the Z-axis of a Cartesian coordinate system), with the amount of precession dependent upon the energy and duration of the RF pulse. A "90.degree." RF pulse is defined as that pulse of RF energy causing the magnetization M to nutate through an angle of 90.degree. from the direction of the B.sub.0 magnetic field, e.g. to move into the X-Y plane (defined by the X-axis and the Y-axis in the Cartesian coordinate system in which the B.sub.0 field is aligned along the Z-axis). Similarly, a "180.degree." RF pulse is defined as that pulse which causes the magnetization M to reverse direction, i.e. move through an angle of 180.degree., from its original direction (e.g. from the positive Z-axis to the negative Z-axis direction). Following the excitation of the nuclei with RF energy, the absorbed energy is re-radiated as an NMR response signal, as the nuclei return to equilibrium. The re-radiated energy is both emitted as radio waves and transferred to molecules, of the sample being investigated, surrounding each re-radiating nucleus.
NMR response signals originating at different spatial locations within the sample can be distinguished by causing their respective resonant frequencies to differ in some predetermined manner. If one or more magnetic field gradients are applied to the sample, and if each gradient filed is of sufficient strength to spread the NMR response signal spectra in a predetermined manner, then each nuclear spin along the direction of at least one of the field gradients experiences magnetic field strength different from the magnetic field strength experienced by other nuclear spins, and therefore resonates at a Larmor frequency different from that of other nuclear spins, as predicted by the Larmor equation. Spatial position of each nucleus contributing to the total NMR response signal can be determined by Fourier analysis, when coupled with knowledge of the applied magnetic field gradient configuration.
The return of nuclear spins to equilibrium, following RF excitation, is referred to as "relaxation". The relaxation process is characterized by two time constants, T.sub.1 and T.sub.2, both of which are measures of motion on the molecular level. The spatial distribution of the T.sub.1 and T.sub.2 constants throughout the sample provides information as to the coupling, or interaction, of the nuclei with their surroundings (T.sub.1) or with each other (T.sub.2) and both provide useful imaging information.
The time constant T.sub.1 is referred to as the "longitudinal", or "spin-lattice", NMR relaxation time, and is a measure of time required for the magnetization M to return to equilibrium; that is, time constant T.sub.1 measures the tendency of the nuclear spins to realign themselves with the total field B.sub.0, after cessation of RF excitation which has moved the spins away from the B.sub.0 field direction. The rate of return to equilibrium is dependent upon how fast the stimulated nuclei can transfer energy to the surrounding sample material, or sample lattice. Time constant T.sub.1, for proton (.sup.1 H) NMR can vary from a few milliseconds in liquids to several minutes or hours in solids. In biological tissue, the typical range of time constant T.sub.1 is from about 30 milliseconds to about 3 seconds.
The time constant T.sub.2 is referred to as the "transverse", or "spin-spin", relaxation time and is a measure of how long the excited nuclear spins oscillate in phase with one another. Immediately after an RF excitation pulse, the nuclear spins are in phase and precess together; however, each nuclear spin behaves like a magnet which generates a magnetic field that affects other spinning nuclei in its vicinity (generating spin-spin interaction). As each nuclear moment experiences its own slightly different magnetic field, due to the spin of adjacent nuclei, that magnetic moment will precess at a different rate and dephase with respect to the other spins, thereby reducing the observable NMR response signal with a time constant T.sub.2. Time constant T.sub.2 can vary from a few microseconds in solids to several seconds in liquids, and is always less than or equal to time constant T.sub.1. In biological tissue, the typical range of time constant T.sub.2, for .sup.1 H NMR, is from about 5 milliseconds to about 3 seconds.
If the static magnetic field B.sub.0 itself has inherent inhomogeneities, as is typically the case with a field generated by a practical magnet, these inherent inhomogeneities produce additional dephasing action, which hastens the decay of the NMR signal. This additional dephasing action occurs because nuclear spins in different spatial locations are exposed to slighly different magnetic field values and therefore resonate at slightly different frequencies. This new relaxation time, which includes the effects of magnetic inhomogeneities, is generally designated T.sub.2 *(T.sub.2 star), where T.sub.2 *.ltoreq.T.sub.2.
In addition to the effect of spin time constants upon the magnitude of the RF energy re-radiated from a particular nuclei, the frequency of the RF electromagnetic energy re-radiated from any particular nuclei can also be affected by local chemical shifts. Chemical shifts occur where the NMR frequency of resonant nuclei, of the same type in a given molecule, differ because of the different magnetic environments, which are themselves produced by differences in the chemical environment of each of the multiplicity of nuclei. This chemical environment difference may occur, for example, due to electrons partly screening the nucleus of a particular atom from the magnetic field; the nucleus therefore has a somewhat-reduced resonant frequency due to the somewhat-reduced total mangnetic field to which that nucleus is subjected. The degree of shielding caused by electrons depends upon the total environment of the nucleus, and therefore the chemical-shift spectrum of a given molecule is unique and can be utilized for identification. Because the resonant frequency (and therefore the absolute chemical shift) is dependent upon the strength of the total applied field, the chemical-shift spectrum is generally expressed as a fractional shift, in parts-per-million (ppm), of the NMR frequency, relative to an arbitrary reference compound. Illustratively, the range of chemical shifts is about 10 ppm for protons (.sup.1 H) to about 200 ppm for carbon (.sup.13 C); other nuclei of interest, such as phosphorous (.sup.31 P) for example, have intermediate chemical shift ranges, e.g. 30 ppm. In order to perform chemical-shift spectroscopy, in which such small chemical shifts are discernible, the homogeneity of the static B.sub.0 magnetic field must be better than the differences in chemical shifts of the spectral peaks to be observed, and is typically required to be much better than one part in 10.sup.6 (1 ppm).
Thus, nuclear magnetic resonance investigation offers two non-invasive probes for detection and diagnosis of disease states in an organic sample: proton (.sup.1 H) magnetic resonance imaging, which can provide images of the internal human anatomy with excellent soft-tissue contrast brought about by the relatively large differences in NMR relaxation times; and localized phosphorous (.sup.31 P) and carbon (.sup.13 C) chemical-shift spectroscopic imaging to provide direct access to metabolic processes for the assessment of damaged tissue and its response to therapy. In addition, the feasibility of imaging naturally-abundance sodium (.sup.23 Na) and artifically-introduced fluorine (.sup.19 F) has recently been demonstrated and may find clinical applications in the near future. It is well known that the magnetic field requirements for .sup.1 H imaging can be met at static field strengths below 0.5 Tesla (T) and that spectroscopy typically requires a magnetic field strength in excess of 1 T, with a much higher degree of homogeneity across the examination region than required for imaging. It is also well known that the signal-to-noise ratio of the NMR signal improves with increasing magnetic field strength, if the rest of the NMR system is optimized. It has been widely speculated, in the literature of the art, that human head and body proton imaging is not feasible above a main field strength of about 0.5 T, owing to the dual problems of RF field penetration into the sample to be investigated and to the difficulty of NMR coil design, at the relatively high NMR frequencies associated with the higher-magnitude static fields. Therefore, by at least implication, a single magnetic resonance system having a single high magnetic field magnitude, in excess of about 0.7 T, for producing proton images and localized chemical-shift spectra from anatomical samples, such as the head, limbs and body of human beings, has been considered experimentally incompatible. A system enabling the performance of both high-field NMR imaging and chemical-shift spectroscopy for medical applications with the human body, and for the analysis distribution of non-ferromagnetic samples (e.g. for analysis and distribution of hydrocarbon deposits and oil-bearing shale sediments, for general morphological and chemical analysis of heterogeneous non-ferromagnetic samples, and the like) is highly desirable. Such a system requires more stringent excitation control and response analysis parameters than a system operating at a lower field or operating in only one of the imaging or spectroscopic regimes. The digital subsystem, having direct control over almost all system variable parameters, must be capable of properly utilizing the magnetic assets of the system to effect superior imaging and spectroscopic analysis.